MRI conditionally safe lead with low-profile multi-layer conductor for longitudinal expansion

CROSS-REFERENCE TO RELATED APPLICATION

This application claims priority to Provisional Application No. 61/291,556, filed Dec. 31, 2009, which is herein incorporated by reference in its entirety.

TECHNICAL FIELD

Various embodiments of the present invention generally relate to implantable medical devices. More specifically, embodiments of the present invention relate to MRI conditionally safe lead conductors.

BACKGROUND

When functioning properly, the human heart maintains its own intrinsic rhythm and is capable of pumping adequate blood throughout the body's circulatory system. However, some individuals have irregular cardiac rhythms, referred to as cardiac arrhythmias, which can result in diminished blood circulation and cardiac output. One manner of treating cardiac arrhythmias includes the use of a pulse generator (PG) such as a pacemaker, an implantable cardioverter defibrillator (ICD), or a cardiac resynchronization (CRT) device. Such devices are typically coupled to a number of conductive leads having one or more electrodes that can be used to deliver pacing therapy and/or electrical shocks to the heart. In atrioventricular (AV) pacing, for example, the leads are usually positioned in a ventricle and atrium of the heart, and are attached via lead terminal pins to a pacemaker or defibrillator which is implanted pectorally or in the abdomen.

Magnetic resonance imaging (MRI) is a non-invasive imaging procedure that utilizes nuclear magnetic resonance techniques to render images within a patient's body. Typically, MRI systems employ the use of a magnetic coil having a magnetic field strength of between about 0.2 to 3 Teslas. During the procedure, the body tissue is briefly exposed to radio frequency (RF) pulses of electromagnetic energy in a plane perpendicular to the magnetic field. The resultant electromagnetic energy from these pulses can be used to image the body tissue by measuring the relaxation properties of the excited atomic nuclei in the tissue. In some cases, imaging a patient's chest area may be clinically advantageous. In a chest MRI procedure, implanted pulse generators and leads may also be exposed to the applied electromagnetic fields.

SUMMARY

The various embodiments of the present invention relate to an implantable electrical lead including a multi-layer coil conductor configured such that the lead exhibits a predetermined axial stiffness or bending stiffness. In Example 1, an implantable electrical lead comprising a flexible body, a connector assembly, an electrode, and a multi-layer coil conductor. The flexible body has a proximal region including a proximal end, a distal region, and a longitudinal body lumen extending from the proximal end through the distal region. The connector assembly is configured to mechanically and electrically connect the implantable electrical lead to an implantable pulse generator. The electrode is coupled to the flexible body in the distal region, and the multi-layer coil conductor extends within the longitudinal body lumen from the connector assembly to at least the electrode. The multi-layer coil conductor includes a first coil layer and a second coil layer disposed about the first coil layer. The first and second coil layers are configured such that the multi-layer coil conductor has an axial stiffness substantially equal to an axial stiffness of the lead body adjacent to the multi-layer coil conductor.

In Example 2, the implantable electrical lead of Example 1, wherein the each of the first and second coil layers is characterized by a filar thickness, a coil outside diameter, and a coil pitch, and wherein one or more of the filar thickness, the coil outside diameter, and the coil pitch are selected such that the axial stiffness of the multi-layer coil conductor is substantially equal to the axial stiffness of the lead body adjacent to the multi-layer coil conductor.

In Example 3, the implantable electrical lead of Examples 1 or 2, wherein a filar material and a coil wind geometry are further selected such that the axial stiffness of the multi-layer coil conductor is substantially equal to the axial stiffness of the lead body adjacent to the multi-layer coil conductor.

In Example 4, the implantable electrical lead of any of Examples 1 through 3, wherein the multi-layer coil conductor is mechanically and electrically connected to a second electrode coupled to the lead body.

In Example 5, the implantable electrical lead of any of Examples 1 through 4, wherein the multi-layer coil conductor allows the second electrode to move axially relative to the terminal connector when under tensile load.

In Example 6, the implantable electrical lead of any of Examples 1 through 5, further comprising a layer of insulating material between the first and second coil layers.

In Example 7, the implantable electrical lead of any of Examples 1 through 5, wherein each of the filars of the first and second coil layers is individually insulated.

In Example 8, an implantable electrical lead, comprising a flexible body, a connector assembly, an electrode, and a multi-layer coil conductor. The flexible body has a proximal region including a proximal end, a distal region, and a longitudinal body lumen extending from the proximal end through the distal region. The connector assembly is configured to mechanically and electrically connect the implantable electrical lead to an implantable pulse generator. The electrode is coupled to the flexible body in the distal region, and the multi-layer coil conductor extends within the longitudinal body lumen from the connector assembly to at least the electrode, and includes a first coil layer a second coil layer disposed about the first coil layer. At least one parameter of the first and the second coil layer is configured such that the lead exhibits a predetermined axial stiffness or bending stiffness.

In Example 9, the implantable electrical lead of Example 8, further comprising a third coil layer disposed around the second coil layer, wherein at least one parameter of the third coil layer is configured such that the lead exhibits a predetermined axial bending stiffness.

In Example 10, the implantable lead of Examples 8 or 9, wherein the at least one parameter of the first and second coil layers is at least one of a coil filar thickness, a number of filars, a filar material, a coil winding direction, or a coil diameter.

In Example 11, the implantable lead of any of Examples 8 through 10, wherein lead of wherein the first and second coil layers are configured such that the multi-layer coil conductor has an axial stiffness substantially equal to an axial stiffness of the lead body adjacent to the multi-layer coil conductor.

In Example 12, the implantable lead of claim any of Examples 8 through 11, wherein each of the coil layers is formed from one or more filars of a silver core DFT wire having an MP35N cladding.

In Example 13, the implantable electrical lead of any of Examples 8 through 12, wherein the multi-layer coil conductor is mechanically and electrically connected to a second electrode coupled to the lead body.

In Example 14, the implantable electrical lead of Example 13, wherein the multi-layer coil conductor allows the second electrode to move axially relative to the terminal connector when under tensile load.

In Example 15, an implantable electrical lead, comprising a flexible body, a connector assembly, an electrode, and a multi-layer coil conductor. The flexible body has a proximal region including a proximal end, a distal region, and a longitudinal body lumen extending from the proximal end through the distal region. The connector assembly is configured to mechanically and electrically connect the implantable electrical lead to an implantable pulse generator. The electrode is coupled to the flexible body in the distal region, and the multi-layer coil conductor extends within the longitudinal body lumen from the connector assembly to at least the electrode, and includes a first coil layer and a second coil layer disposed about the first coil layer. At least one parameter of the first and the second coil layer is configured such that the multi-layer coil conductor has a pre-determined spring constant along at least a portion of its length.

In Example 16, the implantable electrical lead of Example 15 wherein the pre-determined spring constant is sufficient to allow the multi-layer coil conductor to substantially return the lead to its original length upon axial elongation.

In Example 17, the implantable electrical lead of Examples 15 or 16, wherein the multi-layer coil conductor is mechanically and electrically connected to a second electrode coupled to the lead body.

In Example 18, the implantable electrical lead of Examples 15 through 17, wherein the multi-layer coil conductor allows the second electrode to move axially relative to the terminal connector when under tensile load.

BRIEF DESCRIPTION OF THE DRAWINGS

<figref idrefs="DRAWINGS">FIG. 1</figref> is a schematic illustration of a medical system including an MRI scanner, and an implantable cardiac rhythm management system implanted within a torso of a human patient according to various embodiments of the present invention;

<figref idrefs="DRAWINGS">FIG. 2A</figref> is a schematic view of an illustrative pulse generator and lead implanted within the body of a patient which may be used in accordance with some embodiments of the present invention;

<figref idrefs="DRAWINGS">FIG. 2B</figref> is a schematic view showing a simplified equivalence circuit for the lead of <figref idrefs="DRAWINGS">FIG. 2A</figref>

<figref idrefs="DRAWINGS">FIG. 3</figref> is a schematic view illustrating an exemplary lead that may be used in accordance with one or more embodiments of the present invention;

<figref idrefs="DRAWINGS">FIG. 4</figref> is a transverse cross-sectional view of the lead of <figref idrefs="DRAWINGS">FIG. 3</figref> taken along the line <b>4</b>-<b>4</b> in <figref idrefs="DRAWINGS">FIG. 3</figref>; and

<figref idrefs="DRAWINGS">FIG. 5</figref> illustrates invention detail of a multi-layer coil conductor utilized in the lead of <figref idrefs="DRAWINGS">FIG. 3</figref> according to an embodiment of the present invention.

The drawings have not necessarily been drawn to scale. For example, the dimensions of some of the elements in the figures may be expanded or reduced to help improve the understanding of the embodiments of the present invention. While the invention is amenable to various modifications and alternative forms, specific embodiments have been shown by way of example in the drawings and are described in detail below. The intention, however, is not to limit the invention to the particular embodiments described. On the contrary, the invention is intended to cover all modifications, equivalents, and alternatives falling within the scope of the invention as defined by the appended claims.

DETAILED DESCRIPTION

As explained in further detail below, various embodiments of the present invention relate to cardiac rhythm management (CRM) systems incorporating new lead designs advantageously adapted for operation in a magnetic resonance imaging (MRI) environment. In some embodiments, the leads including conductor designs configured to provide suitable electrical performance for delivering pacing and/or defibrillation shock therapy and to minimize the lead's reaction to applied electromagnetic energy during MRI procedures.

In the following description, for the purposes of explanation, numerous specific details are set forth in order to provide a thorough understanding of embodiments of the present invention. It will be apparent, however, to one skilled in the art that embodiments of the present invention may be practiced without some of these specific details.

While, for convenience, some embodiments are described with reference to IMDs in the presence of MRI scanners, embodiments of the present invention may be applicable to various other physiological measurements, treatments, IMD devices, lead types, and other non-invasive examination techniques in which conductive leads are exposed to time varying magnetic fields. As such, the applications discussed herein are not intended to be limiting, but instead exemplary. In addition, various embodiments are applicable to all levels of sensory devices from a single IMD with a sensor to large networks of sensory devices.

<figref idrefs="DRAWINGS">FIG. 1</figref> is a schematic illustration of a medical system <b>100</b> including an MRI scanner <b>110</b>, an implantable cardiac rhythm management (CRM) system <b>115</b> implanted within a torso of a human patient <b>120</b>, and one or more external device(s) <b>130</b> according to various embodiments. The external device(s) <b>130</b> are capable of communicating with the CRM system <b>115</b> implanted within the patient <b>120</b>. In the embodiment shown in <figref idrefs="DRAWINGS">FIG. 1</figref>, the CRM system <b>115</b> includes a pulse generator (PG) <b>140</b> and a lead <b>150</b>. During normal device operation, the PG <b>140</b> is configured to deliver electrical therapeutic stimuli to the patient's heart <b>160</b> for providing tachycardia ventricular fibrillation, anti-bradycardia pacing, anti-tachycardia pacing, cardiac resynchronization therapy, and/or other types of therapy.

Thus, in the illustrated embodiment, the PG <b>140</b> can be an implantable device such as a pacemaker, an ICD, a cardiac resynchronization therapy (CRT) device, a CRT device with defibrillation capabilities (a CRT-D device), or a comparable device. The PG <b>140</b> can be implanted subcutaneously within the body, typically at a location such as in the patient's chest. In some embodiments, PG <b>140</b> can be implanted in or near the abdomen.

The external device(s) <b>130</b> may be a local or remote terminal or other device (e.g., a computing device and/or programming device), operable to communicate with the PG <b>140</b> from a location outside of the patient's body. According to various embodiments, external device <b>130</b> can be any device external to the patient's body that is telemetry enabled and capable of communicating with the PG <b>140</b>. Examples of external devices can include, but are not limited to, programmers (PRM), in-home monitoring devices, personal computers with telemetry devices, MRI scanner with a telemetry device, manufacturing test equipment, or wands. In some embodiments, the PG <b>140</b> communicates with the remote terminal <b>130</b> via a wireless communication interface. Examples of wireless communication interfaces can include, but are not limited to, radio frequency (RF), inductive, and acoustic telemetry interfaces.

<figref idrefs="DRAWINGS">FIG. 2A</figref> is a more detailed schematic view of the CRM system <b>115</b> including the illustrative PG <b>140</b> equipped with the lead <b>150</b> implanted within the body of a patient. In the embodiments depicted, CRM system <b>115</b> includes PG <b>140</b> implanted near the patient's heart <b>160</b> and lead <b>150</b> having a distal portion implanted with the patient's heart <b>160</b>. As can be seen in <figref idrefs="DRAWINGS">FIG. 2A</figref>, the heart <b>160</b> includes a right atrium <b>210</b>, a right ventricle <b>220</b>, a left atrium <b>230</b>, and a left ventricle <b>240</b>.

The lead <b>150</b> has a flexible body <b>200</b> including a proximal region <b>205</b> and a distal region <b>250</b>. As shown, the lead <b>150</b> is coupled to the PG <b>140</b>, and the distal region <b>250</b> of the lead body <b>200</b> is at least partially implanted at a desired location within the right ventricle <b>220</b>. As further shown, the lead <b>150</b> includes a pair of coil electrodes <b>255</b>, <b>257</b> along the distal region <b>250</b>, such that when implanted as shown in <figref idrefs="DRAWINGS">FIG. 2A</figref>, they are positioned within the right ventricle <b>220</b> and right atrium <b>210</b>, respectively. As explained and illustrated in further detail below, the lead <b>150</b> includes one or more electrical conductors within the lead body <b>200</b> (not visible in <figref idrefs="DRAWINGS">FIG. 2A</figref>) electrically coupling the electrodes <b>255</b>, <b>257</b> to circuitry and other electrical components within the PG <b>140</b> for transmitting intrinsic cardiac signals from the heart <b>160</b> to the PG <b>140</b> and also for transmitting electrical shocks or low-voltage pacing stimuli to the heart <b>160</b> via the electrodes <b>255</b>, <b>257</b> or additional electrodes (not shown in <figref idrefs="DRAWINGS">FIG. 2A</figref>).

Although the illustrative embodiment depicts only a single lead <b>150</b> inserted into the patient's heart <b>160</b>, in other embodiments multiple leads can be utilized so as to electrically stimulate other areas of the heart <b>160</b>. In some embodiments, for example, the distal portion of a second lead (not shown) may be implanted in the right atrium <b>210</b>. In addition, another lead may be implanted within the coronary venous system to facilitate pacing the left ventricle, i.e., in a CRT or CRT-D system providing bi-ventricular pacing, as is known in the art. Other types of leads such as epicardial leads may also be utilized in addition to, or in lieu of, the lead <b>150</b> depicted in <figref idrefs="DRAWINGS">FIGS. 1-2</figref>. In short, the various embodiments of the present invention contemplate any multi-lead combinations and configurations for use in CRM systems <b>115</b>, whether now known or later developed.

During operation, the lead <b>150</b> conveys electrical signals between the heart <b>160</b> and the PG <b>140</b>. For example, in those embodiments where the PG <b>140</b> has pacing capabilities, the lead <b>150</b> can be utilized to deliver electrical therapeutic stimulus for pacing the heart <b>160</b>. In those embodiments where the PG <b>140</b> is an ICD, the lead <b>150</b> can be utilized to deliver high voltage electric shocks to the heart <b>160</b> via the electrodes <b>255</b>, <b>257</b> in response to an event such as a ventricular fibrillation.

As explained in detail below, various embodiments of the present invention relate to new lead designs that allow for improved mechanical characteristics and safe operation in an MRI environment. In some embodiments, traditional conductor cables are replaced with low profile multi-layer coil conductors. The multi-layer coil conductors allow lead designers to maintain a small lead profile in pace/sense lead applications while better controlling axial elongation of the lead under tensile load.

<figref idrefs="DRAWINGS">FIG. 2B</figref> is a schematic view showing a simplified equivalence circuit <b>260</b> for the lead <b>150</b> of <figref idrefs="DRAWINGS">FIG. 2A</figref>, representing the RF energy picked up on the lead <b>150</b> from RF electromagnetic energy produced by an MRI scanner. As shown in <figref idrefs="DRAWINGS">FIG. 2B</figref>, voltage (Vi) <b>265</b> in the circuit <b>260</b> represents an equivalent source of energy picked up by the lead <b>150</b> from the MRI scanner. During magnetic resonance imaging, the length of the lead <b>150</b> functions similar to an antenna, receiving the RF energy that is transmitted into the body from the MRI scanner. Voltage (Vi) <b>265</b> in <figref idrefs="DRAWINGS">FIG. 2B</figref> may represent, for example, the resultant voltage received by the lead <b>150</b> from the RF energy. The RF energy picked up by the lead <b>150</b> may result, for example, from the rotating RF magnetic field produced by an MRI scanner, which generates an electric field in the plane perpendicular to the rotating magnetic field vector in conductive tissues. The tangential components of these electric fields along the length of the lead <b>150</b> couple to the lead <b>150</b>. The voltage (Vi) <b>265</b> is thus equal to the integration of the tangential electric field (i.e., the line integral of the electric field) along the length of the lead <b>150</b>.

The Zl parameter <b>270</b> in the circuit <b>260</b> represents the equivalent impedance exhibited by the lead <b>150</b> at the RF frequency of the MRI scanner. The impedance value Zl <b>270</b> may represent, for example, the inductance or the equivalent impedance resulting from the parallel inductance and the coil turn by turn capacitance exhibited by the lead <b>150</b> at an RF frequency of 64 MHz for a 1.5 Tesla MRI scanner, or at an RF frequency of 128 MHz for a 3 Tesla MRI scanner. The impedance Zl of the lead <b>150</b> is a complex quantity having a real part (i.e., resistance) and an imaginary part (i.e., reactance).

Zb <b>275</b> in the circuit <b>260</b> may represent the impedance of the body tissue at the point of lead contact. Zc <b>280</b>, in turn, may represent the capacitive coupling of the lead <b>150</b> to surrounding body tissue along the length of the lead <b>150</b>, which may provide a path for the high frequency current (energy) to leak into the surrounding tissue at the RF frequency of the MRI scanner. Minimizing the absorbed energy (represented by source Vi <b>265</b>) reduces the energy that is transferred to the body tissue at the point of lead contact with the body tissue.

As can be further seen in <figref idrefs="DRAWINGS">FIG. 2B</figref>, the lead <b>150</b> has some amount of leakage into the surrounding tissue at the RF frequency of the MRI scanner. As further indicated by <b>275</b>, there is also an impedance at the point of contact of the lead electrodes <b>255</b>, <b>257</b> to the surrounding body tissue within the heart <b>160</b>. The resulting voltage Vb delivered to the body tissue may be related by the following formula: <br /><i>Vb=Vi Zbe</i>/(<i>Zbe+Zl</i>), where <i>Zbe=Zb </i>in parallel with <i>Zc. </i>

The temperature at the tip of the lead <b>150</b> where contact is typically made to the surrounding tissue is related in part to the power dissipated at <b>275</b> (i.e., at “Zb”), which, in turn, is related to the square of Vb. To minimize temperature rises resulting from the power dissipated at <b>275</b>, it is thus desirable to minimize Vi (<b>265</b>) and Zc (<b>280</b>) while also maximizing the impedance Zl (<b>270</b>) of the lead <b>150</b>. In some embodiments, the impedance Zl (<b>270</b>) of the lead <b>150</b> can be increased at the RF frequency of the MRI scanner, which aids in reducing the energy dissipated into the surrounding body tissue at the point of contact <b>275</b>.

In the various embodiments described in further detail below, the impedance of the lead <b>150</b> can be increased by adding inductance to the lead <b>150</b> and/or by a suitable construction technique. For example, in various embodiments, the inductance of the lead <b>150</b> is increased by selectively configuring the conductors used to supply electrical energy to the electrodes <b>255</b>, <b>257</b>.

<figref idrefs="DRAWINGS">FIG. 3</figref> illustrates in further detail the exemplary lead <b>150</b> that may be used in accordance with one or more embodiments of the present invention. As shown in <figref idrefs="DRAWINGS">FIG. 3</figref>, the lead body <b>200</b> includes a proximal end <b>305</b>, and the lead <b>150</b> further includes a connector assembly <b>310</b> coupled to the proximal end <b>305</b> of the lead body, the coil electrodes <b>255</b>, <b>257</b>, and a tip electrode <b>312</b> which operates in the illustrated embodiment as a pace/sense electrode. Depending on the functional requirements of the PG <b>140</b> (see <figref idrefs="DRAWINGS">FIG. 1</figref>), and the therapeutic needs of the patient, the distal region <b>250</b> of the lead <b>150</b> may include additional electrodes. For example, in some embodiments, the pair of coil electrodes <b>255</b>, <b>257</b> can be used to function as shocking electrodes for providing a defibrillation shock to the heart <b>160</b>. In some embodiments, the lead <b>150</b> can include a low-voltage (e.g., ring) electrode proximal to the distal tip of the lead <b>150</b> which is also operable as a pace/sense electrode, which can be included in addition to, or in lieu of, the tip electrode <b>312</b>. In short, the lead <b>150</b> can incorporate any number of electrode configurations within the scope of the embodiments of the present invention.

In the illustrated embodiment, the connector assembly <b>310</b> includes a connector body <b>320</b> and a terminal pin <b>325</b>. The connector assembly <b>310</b> is coupled to the lead body <b>200</b> and can be configured to mechanically and electrically couple the lead to a header on PG <b>140</b> (see <figref idrefs="DRAWINGS">FIG. 1</figref> and <figref idrefs="DRAWINGS">FIG. 2</figref>). In various embodiments, the terminal pin <b>325</b> extends proximally from the connector body <b>320</b> and in some embodiments is coupled to an inner conductor (not shown in <figref idrefs="DRAWINGS">FIG. 3</figref>) that extends longitudinally through the lead body <b>200</b> to the tip electrode <b>312</b>. In some embodiments, the terminal pin <b>325</b> can include an aperture extending therethrough communicating with a lumen defined by the inner conductor coil in order to accommodate a guide wire or an insertion stylet.

In various embodiments, the tip electrode <b>312</b> is in the form of an electrically active fixation helix at the distal end of the lead <b>150</b>. In some embodiments, the tip electrode <b>312</b> can be an extendable/retractable helix supported by a mechanism to facilitate longitudinal translation of the helix relative to the lead body as the helix is rotated. In those embodiments, the terminal pin <b>325</b> may be rotatable relative to the connector body <b>320</b> and the lead body <b>200</b> such that rotation of the terminal pin <b>325</b> relative to the lead body <b>200</b> causes the inner conductor, and in turn, the helical tip electrode to rotate and translate longitudinally relative to the lead body <b>200</b>. Various mechanisms and techniques for providing extendable/retractable fixation helix assemblies (both electrically active and passive) are known to those of ordinary skill in the art, and need not be described in greater detail here.

The pace/sense electrode (whether a solid tip electrode as described above or an active-fixation helix such as shown in <figref idrefs="DRAWINGS">FIG. 3</figref>) can be made of any suitable electrically conductive material such as Elgiloy, MP35N, tungsten, tantalum, iridium, platinum, titanium, palladium, stainless steel, as well as alloys of any of these materials.

The coil electrodes <b>255</b>, <b>257</b> can take on any configuration suitable for delivering a relatively high-voltage therapeutic shock to the heart for defibrillation therapy. In various embodiments, the coil electrodes <b>255</b>, <b>257</b> can be made from any suitable electrically conductive material such as those discussed in the preceding paragraph. The lead <b>150</b> also includes a conductor (not shown in <figref idrefs="DRAWINGS">FIG. 3</figref>) within the lead body <b>200</b> electrically connecting the coil electrodes <b>255</b>, <b>257</b> to an electrical contact on the connector assembly <b>310</b>, which in turn is configured to electrically couple the coil electrodes <b>255</b>, <b>257</b> to electrical components within the PG <b>140</b>.

In <figref idrefs="DRAWINGS">FIG. 4</figref> is a transverse cross-sectional view of the lead <b>150</b> taken along the line <b>4</b>-<b>4</b> in <figref idrefs="DRAWINGS">FIG. 3</figref>. As shown in <figref idrefs="DRAWINGS">FIG. 4</figref>, the lead body <b>200</b> includes an inner tubular member <b>380</b> and an outer tubular member <b>385</b> disposed over and bonded to the inner tubular member <b>380</b>. The tubular members <b>380</b>, <b>385</b> can be made from any number of flexible, biocompatible insulative materials, including without limitation, polymers such as silicone and polyurethane, and copolymers thereof. As further shown, the inner tubular member <b>380</b> includes a plurality of lumens <b>390</b>, <b>395</b>, <b>400</b>, and conductors <b>410</b>, <b>415</b>, and <b>420</b> are disposed, respectively, in the lumens <b>390</b>, <b>395</b>, and <b>400</b>. Each of the conductors <b>410</b>, <b>415</b> and <b>420</b> extends longitudinally within the respective lumen <b>390</b>, <b>395</b>, <b>400</b>, and is electrically coupled to an electrode (e.g., the electrodes <b>312</b>, <b>255</b>, or <b>257</b> in <figref idrefs="DRAWINGS">FIG. 3</figref>) and also to an electrical contact of the connector assembly <b>310</b>.

In some embodiments, the lead body <b>200</b> does not include separate, coaxial tubular members, but instead, includes only a single tubular member (e.g., the member <b>380</b>) including one or more longitudinal lumens for housing the requisite conductors. For illustrative purposes, the three lumens <b>390</b>, <b>395</b>, <b>400</b> of the inner tubular member <b>380</b> are shown having different diameters. In other embodiments, however, the relative dimensions and/or locations of the lumens <b>390</b>, <b>395</b>, <b>400</b> may vary from that shown. In addition, the inner tubular member <b>380</b> may include a greater or lesser number of lumens, depending on the particular configuration of the lead <b>150</b>. For example, the inner tubular member <b>380</b> may include a greater number of lumens to house additional conductor wires and/or electrode coils within the lead <b>150</b> for supplying current to other shocking coils and/or pace/sense electrodes.

In the illustrated embodiment, the conductor <b>410</b> is a single layer coil conductor such as would be used in conjunction with a conventional, low-voltage pace/sense electrode, e.g., the tip electrode <b>312</b>. In some embodiments, for example, the coil conductor <b>410</b> is configured to have a relatively high impedance when exposed to electromagnetic energy such as that present during an MRI scan. In various such embodiments, the coil conductor <b>410</b> is configured according to the various embodiments described, for example, in U.S. Patent Application Publication No. 2009/0198314, which is incorporated herein by reference in its entirety. The increased impedance aids in reducing the energy dissipated into the surrounding body tissue at or near the lead electrode(s). In various embodiments, the conductor <b>410</b> and the lumen <b>390</b> are omitted.

As explained in further detail below, the conductors <b>415</b> and <b>420</b> are multi-layer conductor assemblies incorporated into the lead <b>150</b> to provide a conditionally-safe MRI-compatible lead design, as well as to provide improved fatigue resistance and other mechanical properties during delivery and under chronic operating conditions. In various embodiments, the multi-layer coil conductors <b>415</b>, <b>420</b> include multiple coil layers having selectively controlled coil properties (e.g., pitch, outside diameter, filar thickness, etc.) to make a highly inductive, highly conductive, small diameter conductor that has suitable mechanical characteristics for a stimulation/sensing lead body.

In the illustrated embodiment, the coil conductor <b>415</b> is a three-layer conductor and the coil conductor <b>420</b> is a two-layer conductor. In various embodiments, multi-layer coil conductors utilizing more than three coil layers can be utilized. In various embodiments, the single layer coil conductor <b>410</b> can be replaced with a multi-layer coil conductor similar or identical to the coil conductor <b>415</b> and/or <b>420</b>.

<figref idrefs="DRAWINGS">FIG. 5</figref> is a more detailed side view of the coil conductor <b>420</b> according to one embodiment of the present invention. As shown in FIG. <b>5</b>, the multi-layer coil conductor <b>420</b> is a three-layer coil conductor <b>420</b> including an outer coil layer <b>422</b>, a middle coil layer <b>425</b> and an inner coil layer <b>430</b>. The outer coil layer <b>422</b> is disposed about the middle coil layer <b>425</b> which is disposed about the inner coil layer <b>430</b>. The outer, middle, and inner coil layers <b>422</b>, <b>425</b>, <b>430</b> are electrically coupled to one another in parallel at least at their the proximal ends (i.e., at or near the connector assembly) as well as at their distal ends (i.e., at the electrode <b>255</b>), so as to provide parallel conductive paths between the connector assembly <b>310</b> and the electrode <b>255</b>. In various embodiments, the parameters of the coil layers <b>422</b>, <b>425</b> and <b>430</b> are configured such that the multi-layer coil conductor <b>420</b> has a maximum DC resistance of about 3.0-3.5 ohms, and a high impedance when exposed to an external alternating magnetic field characterized by frequencies associated with MRI scans. It will be appreciated that the two-layer coil conductor <b>415</b> can be configured in substantially the same manner as the coil conductor <b>420</b> shown in <figref idrefs="DRAWINGS">FIG. 5</figref> while only including two coil conductor layers.

In some embodiments, the outer layer <b>422</b> of the coil conductor <b>420</b> can be wound to coil diameter D_oof less than about 0.013 inches. In some embodiments, the outer layer is close pitched and can have one or more filars having a maximum filar thickness of about 0.004 inches. In accordance with various embodiments, the outer layer may or may not be present depending on the desired application of the lead. For example, in some cases, the outer layer may be used to change the resistance of the multi-layer coil conductor <b>420</b>. In at least one embodiment, the outer layer <b>422</b> may have a filar thickness of from about 0.001 to about 0.004 inches.

The middle layer <b>425</b> of the coil conductor <b>420</b> is close pitched and can be wound to a coil diameter D_mless than the inner diameter of the outer coil layer <b>422</b>. In some embodiments, the middle layer can have one or more filars having a maximum filar thickness of from about 0.0007 to about 0.003 inches.

The inner coil layer <b>430</b> of the coil conductor <b>420</b> is close pitched and can be wound to a coil diameter D, less than the inner diameter of the middle layer <b>425</b>. In some embodiments, the inner layer can have one or more filars having a maximum filar thickness of from about 0.0007 to about 0.001 inches.

Of course, in various other embodiments, multi-layer coil conductor <b>420</b> utilizes different ranges of dimensions and other parameters (e.g., filar count) of the respective coil layers <b>422</b>, <b>425</b>, <b>430</b> depending on the operational needs for the lead <b>150</b>.

In the various embodiments, the filar material can be any suitable material exhibiting the desired electrical and mechanical properties. In one embodiment, the filars of the outer, middle, and/or inner coil layers <b>422</b>, <b>425</b>, <b>430</b> are made of drawn filled tube (DFT) wire including a silver core (about 40%) with an MP35N or tantalum cladding. In another embodiment, the outer, middle, and/or inner coil layers <b>422</b>, <b>425</b>, <b>430</b> are made of DFT wire including a tantalum core with a MP35N cladding. The coil layers <b>422</b>, <b>425</b>, <b>430</b> may be comprised of the same or different materials, and each of the coil layers <b>422</b>, <b>425</b>, <b>430</b> may include different silver fill levels.

In various embodiments, the filars of each layer of the coil conductor <b>420</b> are wound in the same pitch direction. That is, the individual filars of each coil layer are wound to have either a right-hand pitch or a left-hand pitch when viewing the coil along its longitudinal axis.

In various embodiments, one or more of the coil layers <b>422</b>, <b>425</b>, <b>430</b> has a variable coil pitch along its length, which operates to de-tune the coil conductor <b>420</b> to reduce the effect of externally applied electromagnetic radiation (i.e., due to an MRI chest scan).

In various embodiments, the individual coil layers <b>422</b>, <b>425</b> and <b>430</b> can be separately optimized to each exhibit a different inductance, e.g., by modifying the filar thickness, pitch, and/or coil layer diameter, to further modify the performance of the lead <b>150</b> under MRI conditions.

Some embodiments of the present invention include one or more layers of insulation between one or more of the adjacent coil layers <b>422</b>, <b>425</b> or <b>430</b>. Alternatively, or additionally, in various embodiments, the individual filars or the coil layers <b>422</b>, <b>425</b>, and/or <b>430</b> are individually insulated. Any suitable insulation material can be utilized, if desired. Exemplary insulating materials for the filars and/or between the coil layers include ethylene tetrafluoroethylene (ETFE), polytetrafluoroethylene (PTFE), expanded PTFE (ePTFE), silicone, and the copolymers of the foregoing.

The various embodiments of the lead <b>150</b> described above, advantageously minimize induced currents in the lead conductors resulting from exposure to external MRI electromagnetic fields. This is in contrast to conventional lead systems utilizing stranded cable conductors to transmit the shocking currents from the PG to the shocking electrodes. While such cable conductors provide excellent electrical performance for delivering anti-tachycardia therapy, stranded cable conductors also have a low impedance and thus are susceptible to generation of induced currents when exposed to an alternating electromagnetic field such as that present during an MRI scan. Various embodiments of the present invention can result in lead <b>150</b> exhibiting a temperature rise when exposed to MRI radiation of approximately half that experienced utilizing traditional high energy cables which are conventionally utilized to supply energy to high-voltage shocking coils in defibrillation leads. The high impedance conductor configurations for the lead <b>150</b> described above minimize the effects of MRI radiation while still providing suitable electrical performance for use in anti-tachycardia therapy applications.

In addition, In accordance with some embodiments, the design parameters of the multi-layer coil conductors can be tuned to control the mechanical properties, e.g., effective spring constant and thus stiffness of the overall lead assembly, bending stiffness, and the like. Such tuning may enable the user minimize stresses (loads) on adjacent (distal) polymer components, minimize the potential for shear bond failure, facilitate/control the lead bias shape, and minimize the impact of axial length tolerance stack-up during assembly, in both low and high voltage lead applications. For example, in various embodiments, the shocking coils <b>255</b>, <b>257</b> (see <figref idrefs="DRAWINGS">FIGS. 2A and 3</figref>) are omitted, and the lead <b>150</b> includes only one or more low-voltage pace/sense electrodes such as the tip electrode <b>312</b> and additional ring electrodes along the lead <b>150</b>. In such embodiments, multi-layer coil conductors such as the coil conductors <b>425</b> and/or <b>420</b> can be utilized to provide sufficient electrical performance in a low-profile design and also optimal mechanical characteristics such as axial stiffness relative to that of adjacent insulative components such as the lead body <b>200</b>.

In various embodiments, the multi-layer coil conductor <b>415</b> and/or <b>420</b> can be tuned to control the effective spring constant (expressed in force per unit length) of the coil conductor <b>415</b> and/or <b>420</b> so as to reduce or eliminate the difference in axial stiffness between the conductor coil and the adjacent insulating elements (i.e., lead body <b>200</b> components). By controlling and optimizing the coil conductor spring constant/axial stiffness in this way, the overall axial strength of the polymeric lead body components can be maintained, for example, where the coil conductor <b>415</b>, <b>420</b> terminates proximal to the lead body component. Additionally, by reducing or eliminating large differences in the axial stiffness of the conductors (i.e., by utilizing the multi-layer coil conductors <b>415</b> and/or <b>420</b>) and the parallel polymeric lead body components, shear forces between these elements are advantageously reduced. Utilizing the multi-layer coil conductors <b>415</b>, <b>420</b> in lieu of conductor cables also eliminates snaking of the conductor that can result due to the contraction and release of a tensile axial load on the conductor.

Exemplary design parameters that can be varied to tune/optimize the mechanical characteristics discussed above can include, without limitation, the selection of conductor materials, number of layers, wind geometry, pitch, filar diameter, number of filars, and others.

In addition to the coil conductor configurations described above, the various embodiments of the lead <b>150</b> of the present invention may optionally incorporate other features or techniques to minimize the effects of MRI radiation. For example, in some embodiments, shielding may be added to the lead <b>150</b> to further reduce the amount of electromagnetic energy picked up from the lead <b>150</b>. For example, the energy picked up from the shielding can be coupled to the patient's body along the length of the lead <b>150</b>, preventing the energy from coupling to the lead tip. The transfer of intercepted energy by the shielding along the length of the shielding/lead can also be inhibited by dissipating the energy as resistive loss, using resistive material for the shielding construction.

Various modifications and additions can be made to the exemplary embodiments discussed without departing from the scope of the present invention. For example, while the embodiments described above refer to particular features, the scope of this invention also includes embodiments having different combinations of features and embodiments that do not include all of the described features. Accordingly, the scope of the present invention is intended to embrace all such alternatives, modifications, and variations as fall within the scope of the claims, together with all equivalents thereof.